This invention relates generally to a system and method for generating three-dimensional (3-D) images of objects to permit non-destructive inspection of the object in fields such as, for example, medical diagnostics.
The American Cancer Society estimates that in 2001 approximately 192,200 new cases of invasive breast cancer (Stages I-IV) can be diagnosed among women in the United States. Another 46,400 women can be diagnosed with ductal carcinoma in situ, a non-invasive breast cancer. It is has been estimated that over 40,000 deaths can occur from breast cancer in the United States annually. Early detection of breast cancer is vital since early detection has repeatedly been shown to improve the chance of survival. Currently, mammography is a preferred modality for early detection of breast cancer. However, mammography is problematic due to the use of potentially harmful ionizing radiation. Since asymptomatic women are screened repeatedly and the effects of radiation are cumulative, it is recommended that ionizing radiation be avoided. Other limitations include the following: 1) mammography is a two-dimensional (2D) projection modality and is therefore subject to superposition artifacts (i.e. features lying on or near the same line of projection can easily be obscured or made indistinct.); 2) mammography typically cannot differentiate malignant from benign lesions and therefore a subsequent test such as a biopsy is needed; and 3) mammography has a sensitivity of approximately 90% and therefore does not detect an estimated 8-22% of palpable breast cancers. Another modality, echo ultrasound imaging is commonly used as an adjunct to mammography because of its ability to discriminate a cyst from a solid mass. Studies have shown that echo ultrasound, however, has not proven to be an effective screening modality. Screening is the use of a modality to detect disease in an asyptomatic population. Echo ultrasound has a limited field of view, is not reproducible, and produces results that are a balance between depth of imaging (penetration of ultrasound) and image resolution.
Therefore, there is a need for a new, safe and accurate (sensitive and specific) modality. The system of the present invention is a novel three-dimensional (3D) approach to ultrasound computed tomography which can provide such a modality. In 1974, Greenleaf et al. first published a technique called “Ultrasound Computed Tomography” (UCT); unlike echo ultrasound that visualizes tissue interfaces, UCT measures the acoustic properties of the tissue (sound velocity and sound attenuation), and allows a quantitative image to be reconstructed. Greenleaf, J. F., S. A. Johnson, S. L. Lee, G. T. Herman, E. H. Wood, Algebraic Reconstruction of Spatial Distributions of Acoustic Absorption Within Tissue from Their Two-Dimensional Acoustic Projections. Acoustical Holography, 1974, 5: p. 591-603. Success with this modality was limited due to the limited availability of computational technology in the 1970s and Greeleaf et al., U.S. Pat. No. 4,105,018 titled Acoustic Examination, Material Characterization And Imaging Of The Internal Structure Of A Body By Measurement Of The Time-Of-Flight Of Acoustic Energy Therethrough specifically limited its technology to 2D ultrasound, at col. 8, line 47 to col. 9, line 1, stating that “[t]he advantage of cylindrical and circular cylindrical symmetry in ultrasound image formation is related to the basic property of all cylindrical surfaces; namely, that there is a translation or cylindrical axis. This means that if a cylindrical wave is generated it remains a cylindrical wave in a medium of constant index of refraction . . . This is equivalent to saying that in cylindrical symmetry each ray is contained in one and only one plane . . . Thus when using cylindrical waves the coupling of information between adjacent planes perpendicular to the cylinder axis is minimal or small compared to the coupling occurring with spherical waves. This is a great advantage and saves computer time since several small multi-plane problems are much easier to solve in total than one large multiple plane problem.”
Currently, echo ultrasound is routinely used as an adjunct to X-ray mammography to determine the differentiation of simple cysts from solid masses. However, echo ultrasound cannot differentiate malignant and benign masses. Also, false positive X-ray mammograms result in a large numbers of unnecessary biopsies; in the US approximately 75% of the million biopsies performed each year are benign. Thus, a non-invasive, specific, diagnostic modality such as the system of the present invention is needed.
Another use of the system of the present invention is as a screening modality, (to detect almost any lesion) this is the detection function that X-ray mammography is used. However, X-ray mammography misses 8 to 22% of palpable breast cancers. Standard echo ultrasound has not been proven effective for screening asymptomatic patients largely due to its inability to reliably detect microcalcifications. There is significant evidence in the literature that a UCT imager can be very sensitive for lesion detection. There has been great controversy over the starting age and frequency of X-ray mammographic screenings. This controversy arises mainly because of two limitations of mammography. The first is that mammography does not work well in dense breasts, which most young women have. The current recommendation is that most women start screening at age 40. However, 5% of breast cancers occur in women under 40. American Cancer Society, Surveillance Research, 1999. The second controversy is the potential risk of the cumulative effects of ionizing radiation. This worry has some doctors recommending mammographic screenings every two years. Since the most aggressive tumors need detection the earliest, frequent screenings are desirable. Our UCT imager may not be able to detect microcalcifications, but it may still have utility as a screening modality in a select patient population in which mammography is not indicated.
A third potential utility of a 3D imager is for image-guided biopsies and surgical planning. The location, size, and stage of a lesion are parameters that are required for effective treatment planning. Therefore, we feel that the optimal diagnostic strategy for the detection and diagnosis of breast abnormalities is a non-invasive imaging method that is not only highly accurate (both sensitive and specific) but also gives the size and 3D location of any lesion detected.
There are several other non-invasive modalities that may be used for screening and/or diagnosis of breast cancer including ultrasound (echo), Single Photon Emitted Computed Tomography (SPECT), Positron Emitted Tomography (PET) and Magnetic Resonance Imaging (MRI). MRI is very expensive and requires the injection of contrast agents to detect tumors. SPECT and PET are low-resolution modalities and require the injection of ionizing radiation. There are several newer technologies emerging (i.e. acoustical holography, infrared, electrical, optical, and elasticity methods) but none have yet proven to be the definitive methodology.
History of UCT
The allure of UCT for breast imaging is that it offers the potential to quantitatively image tissue properties. Most of the experimental work to develop an UCT imager was performed in the late 70's and early 80's. In spite of the limited technology available to these investigators, they showed promising results. For example, Greenleaf et al. achieved a sensitivity of 100% for palpable lesions with UCT for a small sample population. Greenleaf, J. F., R. C. Bahn, Clinical Imaging with Transmissive Ultrasonic Computerized Tomography. IEEE Transactions on Biomedical Engineering, 1981. BME-28(2): p. 177-185. Greenleaf et al. also showed that by combining the speed-of-sound with the patient's age and a measure of image texture that malignant and benign lesions could be differentiated. Greenleaf, J. F., R. C. Bahn, Clinical Imaging with Transmissive Ultrasonic Computerized Tomography. IEEE Transactions on Biomedical Engineering, 1981. BME-28(2): p. 177-185. Scherzinger et al. showed that by employing discriminant analysis, using combinations of speed-of-sound and attenuation in and around the lesion, one can accurately differentiate tissue types. Scherzinger, A. L., R. A. Belgam, P. A. Carson, C. R. Meyer, J. V. Sutherland, F. L. Bookstein, T. M. Silver, Assesment of Ultrasonic Computed Tomography in Symptomatic Breast Patients by Discriminant Analysis. Ultrasound in Med. and Biol., 1989. 15(1): p. 21-28. In a larger study (n=78), Schreiman et al. showed that a computer-aided diagnosis using UCT had a sensitivity of 82.5% for the diagnosis of a malignancy. Schreiman, J. S., J. J. Gisvold, J. F. Greenleaf, R. C. Bahn, Ultrasound Transmission Computed Tomography of the Breast. Radiology, 1984. 150: p. 523-530.
One of the main problems that these early investigators encountered was that they could not acquire enough projections (at least not quickly enough) to reconstruct an image without reconstruction artifacts. In a review article in 1993, Jones states that early investigators were often hindered due to the limited memory and processor speed of their current computers, which affected both image acquisition and reconstruction. Jones, H. W., Recent Activity in Ultrasonic Tomography. Ultrasonics, 1993. 31(5): p. 353-360. In addition, the length of time required to acquire a full study of the breast was too long to avoid patient motion and the resulting artifacts. This long imaging time was a byproduct of having to mechanically move the transducers to each scan position and the large number of projections required to reduce reconstruction artifacts. Greenleaf et al., using a specially designed UCT imager, took about 5 minutes to image 8 slices (4 slices at a time, each slice was 3 mm thick with a 7 mm gap between slices) in a clinical trial. Christoyianni, I., E. Dermatas, G. Kokkinakis, Fast Detection of Masses in Computer-Aided Mammography. IEEE Signal Processing Magazine, 2000: p. 54-64. In this clinical UCT prototype imager, 60 projections with 200 samples each were acquired, and the image reconstructed into a 128×128 matrix. Azhari et al. claim that the need for a large number of projections (i.e. 201 projections for a 128×128 pixel image) to reduce reconstruction artifacts makes standard UCT impractical for clinical use. Azhari, H., S. Stolarski, Hybrid Ultrasonic Computed Tomography. Computers and Biomedical Research, 1997. 30: p. 35-48. As recently as 1991, Jago and Whittingham, using a linear array to improve speed of acquisition, required approximately 2 minutes to acquire data for a 2D slice and an additional 2 hours to reconstruct a 64 by 64 matrix. Jago, J. R., T. A. Whittingham, Experimental Studies in Transmission Ultrasound Computed Tomography. Phys. Med. Biol., 1991. 36(11): p. 1515-1527. Andre et al. note that after the initial experimental research, most of the work on UCT, through the mid 1990's, was in theoretical reconstructions and not in experimental designs. Andre, M. P., H. S. Janee, P. J. Martin, G. P. Otto, B. A. Spivey, D. A. Palmer, High-Speed Data Acquisition in a Diffraction Tomography System Employing Large-Scale Toroidal Arrays. International Journal of Imaging Systems Technology, 1997. 8(1): p. 137-147. They attributed this trend to limited technologies and speculate that improved instrumentation has led to a renewed interest in UCT.
There are several limitations to UCT which arise from the behavior of sound as it transverses an inhomogeneous media. These include reflection, refraction, and diffraction. There are a number of methods in the literature to correct or account for these effects. Meyer et al. proposed a method to correct for multipath errors using a parametric multipath modeling and estimation technique. Meyer, C. R., T. L. Chenevert, P. L. Carson, A Method for Reducing Multipath Artifacts in Ultrasonic Computed Tomography. J. Acoust. Soc. Am., 1982.72(3): p. 820-823. In a noiseless case, they showed an improvement in attenuation estimates. Pan and Liu proposed methods for correcting refractive errors. Pan, K. M., C. N. Liu, Tomographic Reconstruction of Ultrasonic Attenuation with Correction for Refractive Errors. IBM J. Res. Develop., 1981. 25(1): p. 71-82. They proposed to scan a small area around the straight line-of-sight and then use several different methods (i.e. maximum, sum, or average of the scan area) to measure attenuation. Chenevert et al. explored methods such as cross-correlation and phase-insensitive arrays. Chenevert, T. L., D. I. Bylski, P. L. Carson, P. H. Bland, D. D. Adler, R. M. Schmitt, Ultrasonic Computed Tomography of the Breast. Radiology, 1984. 152: p. 155-159; and Schmitt, R. M., C. R. Meyer, P. Carson, L, T. L. Chenevert, P. H. Bland, Error Reduction in Through Transmission Tomography Using Large Receiving Arrays with Phase-Insensitive Signal Processing. IEEE Transactions on Sonics and Ultrasonics, 1984. SU-31(4): p. 251-258. Cross-correlation minimizes the chance of noise being mistaken as the arrival of the received signal by comparing the signal to a water-path only signal. The use of a phase-insensitive array results in a better attenuation image, by accounting for refraction. Klepper et al. showed that reconstructing an image, where each pixel is the slope of attenuation vs. frequency, minimizes errors due to reflection and refraction. They used a range of frequencies from 3 MHz to 7 MHz and fit a straight line to the data. Klepper, J. R., G. H. Brandenburger, J. W. Mimbs, B. E. Sobel, J. G. Miller, Application of Phase-Insensitive Detection and Frequency-Dependent Measurements to Computed Ultrasonic Attenuation Tomography. IEEE Transactions on Biomedical Engineering, 1981. BME-28(2): p. 186-201. Greenleaf et al. also showed that the intercept of attenuation vs. frequency could be reconstructed. Greenleaf, J. F., R. C. Bahn, Signal Processing Methods for Transmission Ultrasonic Computerized Tomography. Ultrasonics Symposium Proceedings, 1980: p. 966-972. This is an image of reflection by structures larger than the wavelength and is highly correlated with back-scattered information imaged in B-mode scans.
Several investigators explored the use of ray-tracing, ray-linking and iterative reconstructions to generate more accurate images. Improvements in restoring macrostructural geometric proportions have been shown for object inhomogeneities of 5-10% (the breast has inhomogeneities of about 8%). These methods begin with a straight ray assumption to reconstruct an initial speed-of-sound image. Then ray-linking is used to create “new” projections, which are subsequently used to reconstruct a new image. This process is then iterated. Most of the ray-linking methods involve a technique called “shooting”, iteratively searching for the initial angle of the ray from the transmitter that “hits” within some window around the receiver. These methods are computationally expensive and may not be possible to implement in a 3D imaging system.
Norton proposed an alternative method, which involves transforming the ray equation into an implicit integral equation satisfying the boundary conditions. These equations are then solved via successive approximations. Norton also proposed an explicit expression for the ray equation that is correct to the first order for refractive-index perturbations. Norton, S. J., Computing Ray Trajectories Between Two Points: A Solution to the Ray-Linking Problem. Optical Society of America, 1987. 4(10): p. 1919-1922. Andersen proposed an alternative technique based on rebinning of the projection data. Andersen, A. H., A Ray Tracing Approach to Restoration and Resolution Enhancement in Experimental Ultrasound Tomography. Ultrasonic Imaging, 1990. 12: p. 268-291; and Andersen, A. H., Ray Linking for Computed Tomography by Rebinning of Projection Data. J. Acoust. Soc. Am., 1987. 81(4): p. 1190-1192. In this technique, a new radial coordinate system is constructed passing through the center of the image. Rays are then projected from lines passing through the origin. A rebinning process is used and a new image reconstructed. This method is computationally less expensive than the brute force method typically used in ray-linking, a 60% timesavings. Andersen, A. H., A Ray Tracing Approach to Restoration and Resolution Enhancement in Experimental Ultrasound Tomography. Ultrasonic Imaging, 1990. 12: p. 268-291. We can extend the method of rebinning (center-out) to 3D in our reconstructions, detailed in the method of the present invention.
Diffraction Tomogranhy
An alternative to geometrical acoustics for reconstruction is diffraction tomography, which often uses an approximation (Rytov or Born) to the wave equation to reconstruct images. These approximations are only valid in cases of weak scattering. Several investigators have experimented with diffraction imaging.
In order to obtain a linear approximation to the inhomogeneous wave equation, diffraction tomography is often based on the assumption of weak scatters. This assumption is not valid in the human breast due to highly refractive fat layers under the skin. One potential alternative involves the use of higher order approximations to the wave equation. Another alternative is to use iterative methods to solve the wave equation directly. Both of these alternatives are computationally very expensive. For example, the CPU time on a Cray computer was 2.5 hours to reconstruct a 200×200 pixel image from 200 projections. The reconstructed image was very accurate both qualitatively and quantitatively. Manry and Broschat showed that the incorporation of a priori information reduced the computational time by reducing the number of iterations approximately 40%, but the image becomes discritized to 3 grey levels. Manry, C. W. J., S. L. Broschat, Inverse Imaging of the Breast with a Material Classification Technique. J. Acoust. Soc. Am., 1998. 103(3): p. 1538-1546. Lu et al. have recently published a new method that involves the creation of a reconstruction method in a finite form utilizing a formal parameter. Lu, Z.-Q., C.-H. Tan, Z.-Y. Tao, Q. Xue, Acoustical Diffraction Tomography in a Finite Form and Its Computer Simulations. IEEE Transactions on Ultrasonics, Ferroelectrics, and Frequency Control, 2001. 48(4): p. 969-975. This method utilizes an approximation that is much less restrictive than those in the Born and Rytov approximations.
Much of the work in the area of diffraction tomography is also limited by the assumption that the object is being isonofied by a plane wave, which is not feasible in an imager. Sponheim, N., I. Johansen, Experimental Results in Ultrasonic Tomography Using a Filtered Backpropagation Algorithm. Ultrasonic Imaging, 1991. 13: p. 56-70. Sponheim and Johansen have suggested utilizing a reference wave as a first order correction. There is some theoretical work in which non-plane waves are used. Devaney and Beylkin developed a method for utilizing fan beam (spherical or cylindrical) isonofication for diffraction tomography and for the use of arbitrary transmitter and receiver configurations. Devaney, A. J., Generalized Projection-Slice Theorem for Fan Beam Diffraction Tomography. Ultrasonic Imaging, 1985. 7: p. 264-275; and Devaney, A. J., G. Beylkin, Diffraction Tomography Using Arbitrary Transmitter and Receiver Surfaces. Ultrasonic Imaging, 1984. 6: p. 181-193. Witten et al. included the effects of the transmitter beam pattern in their theoretical design of a practical 2D diffraction tomographer. Witten, A., J. Tuggle, R. C. Waag, A Practical Approach to Ultrasonic Imaging Using Diffraction Tomography. J. Acoust. Soc. Am, 1988. 83(4): p. 1645-1652. It is also interesting to note that they claim that any practical ultrasound based imager must use fixed transducers to eliminate errors due to vibrations and acquire data quickly enough to avoid artifacts due to patient motion. Our clinical design meets these requirements.
Fast (approximately 3 sec per slice) 2D diffraction tomography systems with high-resolution (<1 mm in plane, but 10 mm thick slices) and utilization of cylindrical waves have been previously developed. However, the systems result in nonisotropic voxels. Such a system is discussed in Andre, M. P., H. S. Janee, P. J. Martin, G. P. Otto, B. A. Spivey, D. A. Palmer, High-Speed Data Acquisition in a Diffraction Tomography System Employing Large-Scale Toroidal Arrays. International Journal of Imaging Systems Technology, 1997. 8(1): p. 137-147. Others have also experimented with a 2D system resulting in nonisotropic voxels; the system has an in-plane resolution of 0.5 mm; however, the slice thickness is again 10 mm. This system is discussed in Sponheim, N., I. Johansen, Experimental Results in Ultrasonic Tomography Using a Filtered Backpropagation Algorithm. Ultrasonic Imaging, 1991. 13: p. 56-70. Both of these experimental systems utilize first order Bom or Rytov approximations.
Pixel/voxel number and size is also variable. Isotropic voxels (meaning same dimension in all three directions) is a feature of the present invention. The aforementioned systems do not have isotropic voxels: Many image modalities have good in-plane resolution but have thick slices. This creates partial volume error which is blurring of true tissue properties due to averaging of large sections of the tissue into one value that is displayed in an image.
Most diffraction tomography methods reconstruct only in 2D and thus the 3D scattering effect of the breast is a limiting factor of diffraction tomography, and has not previously been addressed in a practical imaging system. A 3D reconstruction algorithm for diffraction tomography utilizing a filtered back-projection algorithm on the Radon transform has recently reported in Anastasio, M. A., X. Pan, Computationally Efficient and Statistically Robust Image Reconstruction in Three-Dimensional Diffraction Tomography. J. Opt. Soc. Am. A, 2000. 17(3): p. 391-400. The method provides reconstruction that reduces to a series of 2D reconstructions over the 3D volume. This reconstruction is based on the Born or Rytov approximations.
Most of the work in diffraction tomography has been theoretical with few actual experimental devices being tested, and none in 3D. Diffraction tomography suffers from the weak scattering assumption, which is often employed, and is violated by strongly refracting fat layers. Note that phase aberration of ultrasound is not a function of breast size as explained in Trahey, G. E., P. D. Freiburger, L. F. Nock, D. C. Sullivan, In Vivo Measurements of Ultrasonic Beam Distortion in the Breast. Ultrasonic Imaging, 1991. 13: p. 71-90. This is suggestive that the major contributor to phase aberration is subcutaneous fat and not the internal structure of the breast. There have been similar findings in that examination of the wavefront amplitude profiles shows coherent interference, indicating refraction as the cause, as is explained in Zhu, Q., B. D. Steinberg, Wavefront Amplitude Distribution in the Female Breast. J. Acoustical Society of America, 1994. 96(1): p. 1-9. In addition, diffraction tomography is more computationally expensive than ray-based UCT and may be limited the discrete implementation of the reconstruction process. Therefore, diffraction-based reconstructions is not preferred for use in the present invention. Rather, in the present invention, use of geometrical acoustics, with ray tracing to correct for the refraction caused by subcutaneous fat it is preferred.
Ultrasound Tissue Characterization
There has been work both in vivo and ex vivo on trying to characterize the ultrasound characteristics of breast tissue. The results of this work suggest that if an imager were accurate, the speed-of-sound and the attenuation-of-sound could be combined with specialized statistical methods to differentiate tissue types in vivo.
It has been shown that the average speed-of-sound in the breast was 1510 m/s for pre-menapausal women and the speed of sound decreased to 1468 m/s in postmenopausal women. Kossoff, G., E. K. Fry, J. Jellins, Average Velocity of Ultrasound in the Human Female Breast. The Journal of the Acoustical Society of America, 1973. 53(6): p. 1730-1736. The difference was attributed to the increase in fat in the breast, post-menopause. Yang et al., using very precise techniques, showed in a very small sampling of excised tissue that the mean speed-of-sound in malignant tissue was 1560 m/s while the surrounding normal tissue had speeds ranging from 1404 to 1450 m/s. Yang, J. N., A. D. Murphy, E. L. Madsen, J. A. Zagzebski, K. W. Gilchrist, G. R. Frank, M. C. Mcdonald, C. A. Millard, A. Faraggi, C. A. Jaramillo, F. R. Gosset, A Mothod for In Vitro Mapping of Ultrasonic Speed and Density in Breast Tissue. Ultrasonic Imaging, 1991. 13: p. 91-109.
Arditi et al. in excised pig mammary tissue showed that insertion loss was linear vs. frequency over the range of 2 to 9 MHz. Arditi, M., P. D. Ecmonds, J. f. Jensen, C. L. Mortensen, W. C. Ross, P. Schattner, D. N. Stephens, W. Vinzant, Apparatus for Ultrasound Tissue Characterization of Excised Specimens. Ultrasonic Imaging, 1991. 13: p. 280-297. Landini et al. showed that the slope of attenuation vs. frequency was able to distinguish malignant lesions with productive fibrosis. Landini, L., R. Sarnelli, F. Squartini, Frequncy-Dependent Attenuation in Breast Tissue Characterization. Ultrasound in Med. &Biol., 1985. 11(4): p. 599-603. Berger et al. have shown that the slope of attenuation vs. frequency is dependant on the genital life of the patient. Berger, G., P. Laugier, J. C. Thalabard, J. Perrin, Global Breast Attenuation: Control Group and Benign Breast Diseases. Ultrasonic Imaging, 1990. 12: p. 47-57. Edmonds et al. showed that in excised breast tissue the speed-of-sound had the best distinguishing power and that the use of Classification And Regression Trees (CART) aids in tissue differentiation. Edmonds, P. D., C. L. Mortensen, J. R. Hill, S. K. Holland, J. F. Jensen, P. Schattner, A. D. Valdes, Ultrasound Tissue Characterization of Breast Biopsy Specimens. Ultrasonic Imaging, 1991. 13: p. 162-185.
Scherzinger et al., Greenleaf et al., Glover, and Schreiman et al. have all had some success in discriminating tissue types in vivo using 2D UCT and often employing computer-assisted classifications, in spite of the fact that there is overlap of ultrasound properties between tissue types. Glover, G. H., Computerized Time-of-Flight Ultrasonic Tomography for Breast Examination. Ultrasound Med. Biol., 1977. 3: p. 117-127; Scherzinger, A. L., R. A. Belgam, P. A. Carson, C. R. Meyer, J. V. Sutherland, F. L. Bookstein, T. M. Silver, Assesment of Ultrasonic Computed Tomography in Symptomatic Breast Patients by Discriminant Analysis. Ultrasound in Med. and Biol., 1989. 15(1): p. 21-28; Schreiman, J. S., J. J. Gisvold, J. F. Greenleaf, R. C. Bahn, Ultrasound Transmission Computed Tomography of the Breast. Radiology, 1984. 150: p. 523-530.
Because there is overlap between the speed-of-sound and attenuation for normal and malignant tissue, the use of other parameters may be needed to differentiate tissue. These may include: backscatter coefficient, the acoustic nonlinearity parameter (B/A), temperature dependence of the speed-of-sound and attenuation, quantification of the anisotropy of the ultrasonic tissue properties, as well as, various combinations of all these parameters. The design of the present invention allows the acquisition of these parameters with only slight modifications.
Limitations of Early 2D UCT and Proposed Solutions
Although early 2D UCT imagers showed promise for use as an adjunct diagnostic exam, they were not accepted clinically for several reasons. The first was the technical limitations which 1) limited the number of projections that could be acquired in a reasonable time (created reconstruction artifacts), 2) created long reconstruction times, prohibiting the extension of methods to 3D, and 3) limited computer analysis of the images. Our design overcomes these problems by acquiring a 3D image in approximately 120 see, uses current computational power to reconstruct a 3D image, and creates a digital image, which allows for easy implementation of computerized image analysis.
The second limitation is that the effects of refraction and diffraction are 3D phenomena and have previously only been addressed in 2D. Again, one of our goals is to create a 3D UCT imager specifically to correct for 3D refraction.
The present invention overcomes the disadvantages of prior imaging modalities by providing in one embodiment a 3D UCT imager using a cylindrical array, of small piezoelectric elements acting as both transmitters and receivers. This arrangement allows for quick collection of 3D projections (preferably in a cone beam fashion). Specifically, projections can be created between any pair of piezoelectric elements that are lining the image chamber. This geometry creates a cone beam acquisition. In 2D imaging, fan beam acquisition refers to the case where the rays spread from a source like a fan. This configuration creates non-parallel projections. Cone beam is an extension of fan beam acquisition to the 3D case. Thus, the rays spread from a source to create a set of projections shaped like a cone. By utilizing a hemispherical transmission, the present invention, in effect, is a special case of cone beam acquisition, i.e. the acquisition scheme utilizes non-parallel projections. In typical cone beam acquisition, like the configuration used in X-ray computed tomography, a 2D flat surface of receivers is used. In the configuration of at least one embodiment of the present invention, the receivers are on a cylinder. 3D reconstruction will result in true 3D images. The advantages of the 3D UCT imaging system of the present invention are numerous and include: 1) the absence of ionizing radiation, 2) the ability to provide true 3D acquisition, reconstruction, and display, 3) the ability to quantify tissue properties, and thus, differentiate malignant and benign tissue to avoid the need for an invasive biopsy, 4) the ability to image dense breast tissue typically found in young women (i.e. women age 40 or less); 5) the availability of the system for use in frequent follow-up imaging for determination of efficacy of treatment because of the absence of ionizing radiation; and 6) the ability to provide a comparatively comfortable modality, as there is no need for compression of breast tissue when using the system. In addition to the stated advantages, the image created by the system of the present invention is created in a digital format, and has advantages typical of digital image formats, including the use of the 3D digital image with computer aided diagnosis and telemedicine.